Assisted ventilation to match patient respiratory need

ABSTRACT

The apparatus provides for the determination of the instantaneous phase in the respiratory cycle, subject&#39;s average respiration rate and the provision of ventilatory assistance. A microprocessor ( 16 ) receives an airflow signal from a pressure transducer ( 18 ) coupled to a port ( 17 ) at a mask ( 11 ). The microprocessor ( 16 ) controls a servo ( 19 ), that in turn controls the fan motor ( 20 ) and thus the pressure of air delivered by the blower ( 10 ). The blower ( 10 ) is coupled to a subject&#39;s mask (ii) by a conduit ( 12 ). The invention seeks to address the following goals: while the subject is awake and making substantial efforts the delivered assistance should be closely matched in phase with the subject&#39;s efforts; the machine should automatically adjust the degree of assistance to maintain at least a specified minimum ventilation without relying on the integrity of the subject&#39;s chemoreflexes; and it should continue to work correctly in the pesence of large leaks.

FIELD OF THE INVENTION

The invention relates to methods and apparatus for the provision ofventilatory assistance matched to a subject's respiratory need. Theventilatory assistance can be for a subject who is either spontaneouslyor non-spontaneously breathing, or moves between these breathing states.The invention is especially suitable for, but not limited to,spontaneously breathing human subjects requiring longterm ventilatoryassistance, particularly during sleep.

BACKGROUND OF THE INVENTION

Subjects with severe lung disease, chest wall disease, neuromusculardisease, or diseases of respiratory control may require in-hospitalmechanical ventilatory assistance, followed by longterm home mechanicalventilatory assistance, particularly during sleep. The ventilatordelivers air or air enriched with oxygen to the subject, via aninterface such as a nosemask, at a pressure that is higher duringinspiration and lower during expiration.

In the awake state, and while waiting to go to sleep, the subject'sventilatory pattern is variable in rate and depth. Most knownventilatory devices do not accurately match the amplitude and phase ofmask pressure to the subject's spontaneous efforts, leading todiscomfort or panic. Larger amounts of asynchrony also reduce theefficiency of the device. During sleep, there are changes in the neuralcontrol of breathing as well as the mechanics of the subject's airways,respiratory muscles and chest wall, leading to a need for substantiallyincreased ventilatory support. Therefore, unless the device canautomatically adjust the degree of support, the amplitude of deliveredpressure will either be inadequate during sleep, or must be excessive inthe awake state. This is particularly important in subjects withabnormalities of respiratory control, for example centralhypoventilation syndromes, such as Obesity Hypoventilation Syndrome,where there is inadequate chemoreceptor drive, or Cheyne Stokesbreathing such as in patients with severe cardiac failure or after astroke, where there is excessive or unstable chemoreceptor drive.

Furthermore, during sleep there are inevitably large leaks between maskand subject, or at the subject's mouth if this is left free. Such leaksworsen the error in matching the phase and magnitude of the machine'seffort to the subject's needs, and, in the case of mouth leak, reducethe effectiveness of the ventilatory support.

Ideally a ventilatory assistance device should simultaneously addressthe following goals:

(i) While the subject is awake and making substantial ventilatoryefforts, the delivered assistance should be closely matched in phasewith the patient's efforts.

(ii) The machine should automatically adjust the degree of assistance tomaintain at least a specified minimum ventilation, without relying onthe integrity of the subject's chemoreflexes.

(iii) It should continue to work correctly in the presence of largeleaks.

Most simple home ventilators either deliver a fixed volume, or cyclebetween two fixed pressures. They do so either at a fixed rate, or aretriggered by the patient's spontaneous efforts, or both. All such simpledevices fail to meet goal (ii) of adjusting the degree of assistance tomaintain at least a given ventilation. They also largely fail to meetgoal (i) of closely matching the subjects respiratory phase: timeddevices make no attempt to synchronize with the subject's efforts;triggered devices attempt to synchronize the start and end of the breathwith the subject's efforts, but make no attempt to tailor theinstantaneous pressure during a breath to the subject's efforts.Furthermore, the triggering tends to fail in the presence of leaks, thusfailing goal (iii).

The broad family of servo-ventilators known for at least 20 yearsmeasure ventilation and adjust the degree of assistance to maintainventilation at or above a specified level, thus meeting goal (ii), butthey still fail to meet goal (i) of closely matching the phase of thesubject's spontaneous efforts, for the reasons given above. No attemptis made to meet goal (iii).

Proportional assistist ventilation (PAV), as taught by Dr Magdy Younes,for example in Principles and Practice of Mechanical Ventilation,chapter 15, aims to tailor the pressure vs time profile within a breathto partially or completely unload the subject's resistive and elasticwork, while minimizing the airway pressure required to achieve thedesired ventilation. During the inspiratory half-cycle, the administeredpressure takes the form:P(t)=P ₀ +R.f _(RESP)(t)+E.V(t)where R is a percentage of the resistance of the airway, f_(RESP)(t) isthe instantaneous respiratory airflow at time t, E is a percentage ofthe elastance of lung and chest wall, and V(t) is the volume inspiredsince the start of inspiration to the present moment. During theexpiratory half-cycle, V(t) is taken as zero, to produce passiveexpiration.

An advantage of proportional assist ventilation during spontaneousbreathing is that the degree of assistance is automatically adjusted tosuit the subject's immediate needs and their pattern of breathing, andis therefore comfortable in the spontaneously breathing subject.However, there are at least two important disadvantages. Firstly, V(t)is calculated as the integral of flow with respect to time since thestart of inspiration. A disadvantage of calculating V(t) in this way isthat, in the presence of leaks, the integral of the flow through theleak will be included in V(t), resulting in an overestimation of V(t),in turn resulting in a runaway increase in the administered pressure.This can be distressing to the subject. Secondly, PAV relies on thesubject's chemoreceptor reflexes to monitor the composition of thearterial blood, and thereby set the level of spontaneous effort. The PAVdevice then amplifies this spontaneous effort. In subjects with abnormalchemoreceptor reflexes, the spontaneous efforts may either ceaseentirely, or become unrelated to the composition of the arterial blood,and amplification of these efforts will yield inadequate ventilation. Inpatients with existing Cheyne Stokes breathing during sleep, PAV will bydesign amplify the subject's waxing and waning breathing efforts, andactually make matters worse by exaggerating the disturbance. Thus PAVsubstantially meets goal (i) of providing assistance in phase with thesubject's spontaneous ventilation, but cannot meet goal (ii) ofadjusting the depth of assistance if the subject has inadequatechemoreflexes, and does not satisfactorily meet goal (iii).

Thus there are known devices that meet each of the above goals, butthere is no device that meets all the goals simultaneously.Additionally, it is desirable to provide improvements over the prior artdirected to any one of the stated goals.

Therefore, the present invention seeks to achieve, at least partially,one or more of the following:

(i) to match the phase and degree of assistance to the subject'sspontaneous efforts when ventilation is well above a target ventilation,

(ii) to automatically adjust the degree of assistance to maintain atleast a specified minimum average ventilation without relying on theintegrity of the subject's chemoreflexes and to damp out instabilitiesin the spontaneous ventilatory efforts, such as Cheyne Stokes breathing.

(iii) to provide some immunity to the effects of sudden leaks.

DISCLOSURE OF THE INVENTION

In what follows, a fuzzy membership function is taken as returning avalue between zero and unity, fuzzy intersection A AND B is the smallerof A and B, fuzzy union A OR B is the larger of A and B, and fuzzynegation NOT A is 1−A.

The invention discloses the determination of the instantaneous phase inthe respiratory cycle as a continuous variable.

The invention further discloses a method for calculating theinstantaneous phase in the respiratory cycle including at least thesteps of determining that if the instantaneous airflow is small andincreasing fast, then it is close to start of inspiration, if theinstantaneous airflow is large and steady, then it is close tomid-inspiration, if the instantaneous airflow is small and decreasingfast, then it is close to mid-expiration, if the instantaneous airflowis zero and steady, then it is during an end-expiratory pause, andairflow conditions intermediate between the above are associated withcorrespondingly intermediate phases.

The invention further discloses a method for determining theinstantaneous phase in the respiratory cycle as a continuous variablefrom 0 to 1 revolution, the method comprising the steps of:

-   -   selecting at least two identifiable features F_(N) of a        prototype flow-vs-time waveform f(t) similar to an expected        respiratory flow-vs-time waveform, and for each said feature:    -   determining by inspection the phase φ_(N) in the respiratory        cycle for said feature, assigning a weight W_(N) to said phase,    -   defining a “magnitude” fuzzy set M_(N) whose membership function        is a function of respiratory airflow, and a “rate of change”        fuzzy set C_(N), whose membership function is a function of the        time derivative of respiratory airflow, chosen such that the        fuzzy intersection M_(N) AND C_(N) will be larger for points on        the generalized prototype respiratory waveform whose phase is        closer to the said feature F_(N) than for points closer to all        other selected features,    -   setting the fuzzy inference rule R_(N) for the selected feature        F_(N) to be: If flow is M_(N) and rate of change of flow is        C_(N) then phase=φ_(N), with weight W_(N).    -   measuring leak-corrected respiratory airflow,    -   for each feature F_(N) calculating fuzzy membership in fuzzy        sets M_(N) and C_(N),    -   for each feature F_(N) applying fuzzy inference rule R_(N) to        determine the fuzzy extent Y_(N)=M_(N) AND C_(N) to which the        phase is φ_(N), and    -   applying a defuzzification procedure using Y_(N) at phases φ_(N)        and weights W_(N) to determine the instantaneous phase φ.

Preferably, the identifiable features include zero crossings, peaks,inflection points or plateaus of the prototype flow-vs-time waveform.Furthermore, said weights can be unity, or chosen to reflect theanticipated reliability of deduction of the particular feature.

The invention further discloses a method for calculating instantaneousphase in the respiratory cycle as a continuous variable, as describedabove, in which the step of calculating respiratory airflow includes alow pass filtering step to reduce non-respiratory noise, in which thetime constant of the low pass filter is an increasing function of anestimate of the length of the respiratory cycle.

The invention further discloses a method for measuring the instantaneousphase in the respiratory cycle as a continuous variable as describedabove, in which the defuzzification step includes a correction for anyphase delay introduced in the step of low pass filtering respiratoryairflow.

The invention further discloses a method for measuring the averagerespiratory rate, comprising the steps of:

measuring leak-corrected respiratory airflow,

from the respiratory airflow, calculating the instantaneous phase φ inthe respiratory cycle as a continuous variable from 0 to 1 revolution,calculating the instantaneous rate of change of phase dφ/dt, and

calculating the average respiratory rate by low pass filtering saidinstantaneous rate of change of phase dφ/dt.

Preferably, the instantaneous phase is calculated by the methodsdescribed above.

The invention further discloses a method for providing ventilatoryassistance in a spontaneously breathing subject, comprising the steps,performed at repeated sampling intervals, of:

ascribing a desired waveform template function Π(φ), with domain 0 to 1revolution and range 0 to 1,

calculating the instantaneous phase φ in the respiratory cycle as acontinuous variable from 0 to 1 revolution,

selecting a desired pressure modulation amplitude A,

calculating a desired instantaneous delivery pressure as an endexpiratory pressure plus the desired pressure modulation amplitude Amultiplied by the value of the waveform template function Π(φ) at thesaid calculated phase φ, and

setting delivered pressure to subject to the desired delivery pressure.

The invention further discloses a method for providing ventilatoryassistance in a spontaneously breathing subject as described above, inwhich the step of selecting a desired pressure modulation amplitude is afixed amplitude.

The invention further discloses a method for providing ventilatoryassistance in a spontaneously breathing subject as described above, inwhich the step of selecting a desired pressure modulation amplitude inwhich said amplitude is equal to an elastance multiplied by an estimateof the subject's tidal volume.

The invention further discloses a method for providing ventilatoryassistance in a spontaneously breathing subject as described above, inwhich the step of selecting a desired pressure modulation amplitudecomprises the substeps of:

-   -   specifying a typical respiratory rate giving a typical cycle        time,    -   specifying a preset pressure modulation amplitude to apply at        said typical respiratory rate,    -   calculating the observed respiratory rate giving an observed        cycle time, and    -   calculating the desired amplitude of pressure modulation as said        preset pressure modulation amplitude multiplied by said observed        cycle time divided by the said specified cycle time.

The invention further discloses a method for providing ventilatoryassistance in a spontaneously breathing subject, including at least thestep of determining the extent that the subject is adequatelyventilated, to said extent the phase in the respiratory cycle isdetermined from the subject's respiratory airflow, but to the extentthat the subject's ventilation is inadequate, the phase in therespiratory cycle is assumed to increase at a pre-set rate, and settingmask pressure as a function of said phase.

The invention further discloses a method for providing ventilatoryassistance in a spontaneously breathing subject, comprising the stepsof: measuring respiratory airflow, determining the extent to which theinstantaneous phase in the respiratory cycle can be determined from saidairflow, to said extent determining said phase from said airflow but tothe extent that the phase in the respiratory cycle cannot be accuratelydetermined, the phase is assumed to increase at a preset rate, anddelivering pressure as a function of said phase.

The invention further discloses a method for calculating theinstantaneous inspired volume of a subject, operable substantiallywithout run-away under conditions of suddenly changing leak, the methodcomprising the steps of:

determining respiratory airflow approximately corrected for leak,

calculating an index J varying from 0 to 1 equal to the fuzzy extent towhich said corrected respiratory airflow is large positive for longerthan expected, or large negative for longer than expected,

identifying the start of inspiration, and

calculating the instantaneous inspired volume as the integral of saidcorrected respiratory airflow multiplied by the fuzzy negation of saidindex J with respect to time, from start of inspiration.

The invention further discloses a method “A” for providing ventilatoryassistance in a spontaneously breathing subject, the method comprisingthe steps, performed at repeated sampling intervals, of:

determining respiratory airflow approximately corrected for leak,

calculating an index J varying from 0 to 1 equal to the fuzzy extent towhich said respiratory airflow is large positive for longer thanexpected, or large negative for longer than expected,

calculating a modified airflow equal to said respiratory airflowmultiplied by the fuzzy negation of said index J,

identifying the phase in the respiratory cycle,

calculating the instantaneous inspired volume as the integral of saidmodified airflow with respect to time, with the integral held at zeroduring the expiratory portion of the respiratory cycle,

calculating a desired instantaneous delivery pressure as a function atleast of the said instantaneous inspired volume, and

setting delivered pressure to subject to the desired delivery pressure.

The invention further discloses a method “B” for providing ventilatoryassistance in a spontaneously breathing subject, comprising the stepsof:

determining respiratory airflow approximately corrected for leak,

calculating an index J varying from 0 to 1 equal to the fuzzy extent towhich the respiratory airflow is large positive for longer thanexpected, or large negative for longer than expected,

identifying the phase in the respiratory cycle,

calculating a modified respiratory airflow equal to the respiratoryairflow multiplied by the fuzzy negation of said index J,

calculating the instantaneous inspired volume as the integral of themodified airflow with respect to time, with the integral held at zeroduring the expiratory portion of the respiratory cycle,

calculating the desired instantaneous delivery pressure as an expiratorypressure plus a resistance multiplied by the instantaneous respiratoryairflow plus a nonlinear resistance multiplied by the respiratoryairflow multiplied by the absolute value of the respiratory airflow plusan elastance multiplied by the said adjusted instantaneous inspiredvolume, and

setting delivered pressure to subject to the desired delivery pressure.

The invention yet further discloses a method “C” for providing assistedventilation to match the subject's need, comprising the steps of:

describing a desired waveform template function Π(φ), with domain 0 to 1revolution and range 0 to 1,

determining respiratory airflow approximately corrected for leak,

calculating an index J varying from 0 to 1 equal to the fuzzy extent towhich the respiratory airflow is large positive for longer thanexpected, or large negative for longer than expected,

calculating J_(PEAK) equal to the recent peak of the index J,

calculating the instantaneous phase in the respiratory cycle,

calculating a desired amplitude of pressure modulation, chosen toservo-control the degree of ventilation to at least exceed a specifiedventilation,

calculating a desired delivery pressure as an end expiratory pressureplus the calculated pressure modulation amplitude A multiplied by thevalue of the waveform template function Π(φ) at the said calculatedphase φ, and

setting delivered pressure to subject to said desired instantaneousdelivered pressure.

The invention yet further discloses a method for providing assistedventilation to match the subject's need, as described above, in whichthe step of calculating a desired amplitude of pressure modulation,chosen to servo-control the degree of ventilation to at least exceed aspecified ventilation, comprises the steps of:

calculating a target airflow equal to twice the target ventilationdivided by the target respiratory rate,

deriving an error term equal to the absolute value of the instantaneouslow pass filtered respiratory airflow minus the target airflow, and

calculating the amplitude of pressure modulation as the integral of theerror term multiplied by a gain, with the integral clipped to liebetween zero and a maximum.

The invention yet further discloses a method for providing assistedventilation to match the subject's need, as described above, in whichthe step of calculating a desired amplitude of pressure modulation,chosen to servo-control the degree of ventilation to at least exceed aspecified ventilation, comprises the following steps:

-   -   calculating a target airflow equal to twice the target        ventilation divided by the target respiratory rate,    -   deriving an error term equal to the absolute value of the        instantaneous low pass filtered respiratory airflow minus the        target airflow,    -   calculating an uncorrected amplitude of pressure modulation as        the integral of the error term multiplied by a gain, with the        integral clipped to lie between zero and a maximum,    -   calculating the recent average of said amplitude as the low pass        filtered amplitude, with a time constant of several times the        length of a respiratory cycle, and    -   setting the actual amplitude of pressure modulation to equal the        said low pass filtered amplitude multiplied by the recent peak        jamming index J_(PEAK) plus the uncorrected amplitude multiplied        by the fuzzy negation of J_(PEAK).

The invention yet further discloses a method for providing assistedventilation to match the subject's need, and with particular applicationto subjects with varying respiratory mechanics, insufficient respiratorydrive, abnormal chemoreceptor reflexes, hypoventilation syndromes, orCheyne Stokes breathing, combined with the advantages of proportionalassist ventilation adjusted for sudden changes in leak, comprising thesteps, performed at repeated sampling intervals, of:

-   -   calculating the instantaneous mask pressure as described for        methods “A” or “B” above,    -   calculating the instantaneous mask pressure as described for        method “C” above,    -   calculating a weighted average of the above two pressures, and    -   setting the mask pressure to the said weighted average.

The invention yet further discloses apparatus to give effect to each oneof the methods defined, including one or more transducers to measureflow and/or pressure, processor means to perform calculations andprocedures, flow generators for the supply of breathable gas at apressure above atmospheric pressure and gas delivery means to deliverthe breathable gas to a subject's airways.

The apparatus can include ventilators, ventilatory assist devices, andCPAP devices including constant level, bi-level or autosetting leveldevices.

It is to be understood that while the algorithms embodying the inventionare explained in terms of fuzzy logic, approximations to thesealgorithms can be constructed without the use of the fuzzy logicformalism.

BRIEF DESCRIPTION OF THE DRAWINGS

A number of embodiments will now be described with reference to theaccompanying drawings in which:

FIGS. 1 a and 1 b show apparatus for first and second embodiments of theinvention respectively;

FIG. 2 is a pressure waveform function Π(Φ) used in the calculation ofthe desired instantaneous delivery pressure as a function of theinstantaneous phase Φ in the respiratory cycle for a first embodiment ofthe invention;

FIG. 3 shows fuzzy membership functions for calculating the degree ofmembership in each of five magnitude fuzzy sets (“large negative”,“small negative”, “zero”, “small positive”, and “large positive”) fromthe normalized respiratory airflow according to the first embodiment ofthe invention; and

FIG. 4 shows fuzzy membership functions for calculating the degree ofmembership in each of five rate of change fuzzy sets (“rising fast”,“rising slowly”, “steady”, “falling slowly”, and “falling fast”) fromthe normalized rate of change of airflow according to the firstembodiment of the invention;

FIG. 5 is a pressure waveform function Π(Φ) used in the calculation ofthe desired instantaneous delivery pressure as a function of theinstantaneous phase Φ in the respiratory cycle for a second embodimentof the invention;

FIG. 6 shows calculation of a quantity “lead-in” as a function of timesince the most recent mask off-on transition;

FIG. 7 shows a fuzzy membership function for fuzzy set A_(I) as afunction of time since the most recent expiratory-to-inspiratory(negative-to-positive) zero crossing of the respiratory airflow signal,such that the membership function measures the extent to which therespiratory airflow has been positive for longer than expected;

FIG. 8 shows a membership function for fuzzy set B_(I) as a function ofrespiratory airflow, such that the membership function measures theextent to which respiratory airflow is large positive;

FIG. 9 shows an electrical analog of the calculation of a recent peakjamming index J_(PEAK) from the instantaneous jamming index J;

FIG. 10 shows the calculation of the time constant τ used in low passfiltering steps in the calculation of the conductance of a leak, as afunction of the recent peak jamming index J_(PEAK).

FIG. 11 shows a prototypical respiratory flow-time curve, with time onthe x-axis, marking nine features;

FIG. 12 shows membership functions for fuzzy sets “large negative”,“small negative”, “zero”, “small positive”, and “large positive” asfunctions of normalized respiratory airflow according to a secondembodiment of the invention;

FIG. 13 shows membership functions for fuzzy sets “falling”, “steady”,and “rising” as functions of normalized rate of change of respiratoryairflow df/dt according to a second embodiment of the invention;

FIG. 14 shows the membership function for fuzzy set “hypopnea”;

FIG. 15 shows the calculation of the time constant τ for calculation ofnormalized recent ventilation, as a function of “servo gain” being thegain used for servo-control of minute ventilation to at least exceed aspecified target ventilation;

FIG. 16 shows the membership function for fuzzy set “hyperpnea” as afunction of normalized recent ventilation;

FIG. 17 shows the membership function for fuzzy set “big leak” as afunction of leak;

FIG. 18 shows the membership functions for fuzzy sets “switch negative”and “switch positive” as a function of nomalized respiratory airflow;

FIG. 19 shows the membership functions for fuzzy sets “insp_phase” and“exp_phase” as functions of the instantaneous phase in the respiratorycycle φ;

FIG. 20 shows schematically how function W(y), used in defuzzification,calculates the area (shaded) of an isosceles triangle of unit base andheight cut off below height y;

FIGS. 21-26 show actual 60 second flow and pressure tracings from thesecond embodiment of the invention during operation; the vertical scalefor flow (heavy trace) is ±1 L/sec, inspiration upwards and the verticalscale for the pressure (light trace) is 0-25 cmH₂O; where:

FIG. 21 shows that a short central apnea (b) is permitted when effortceases at point (c) after a preceding deep breath (a);

FIG. 22 shows that a central apnea is not permitted when effort ceasesat arrow (a) without a preceeding deep breath;

FIG. 23 is recorded with servo gain set high, and shows that a centralapnea is no longer permitted when effort ceases at arrow (a) despitepreceding deep breathing;

FIG. 24 shows automatically increasing end-inspiratory pressure as thesubject makes voluntarily deeper inspiratory efforts;

FIG. 25 is recorded with a somewhat more square waveform selected, andshows automatically increasing pressure support when the subjectvoluntarily attempts to resist by stiffening the chest wall at point(a);

FIG. 26 shows that with sudden onset of a sever 1.4 L/sec leak at (a),the flow signal returns to baseline (b) within the span of a singlebreath, and pressure continues to cycle correctly throughout; and

FIG. 27 shows an actual 60 second tracing showing respiratory airflow(heavy trace, ±1 L/sec full scale) and instantaneous phase (light trace,0-1 revolution full scale).

DESCRIPTION OF PREFERRED EMBODIMENTS

The two embodiments to be described are ventilators that operate in amanner that seeks to simultaneously achieve the three goals statedabove.

First Embodiment

Apparatus to give effect to a first embodiment of the apparatus is shownin FIG. 1 a. A blower 10 supplies a breathable gas to mask 11 incommunication with the subject's airway via a delivery tube 12 andexhausted via a exhaust diffuser 13. Airflow to the mask 11 is measuredusing a pneumotachograph 14 and a differential pressure transducer 15.The mask flow signal from the transducer 15 is then sampled by amicroprocessor 16. Mask pressure is measured at the port 17 using apressure transducer 18. The pressure signal from the transducer 18 isthen sampled by the microprocessor 16. The microprocessor 16 sends aninstantaneous mask pressure request signal to the servo 19, whichcompares said pressure request signal with actual pressure signal fromthe transducer 18 to the control fan motor 20. The microprocessorsettings can be adjusted via a serial port 21.

It is to be understood that the mask could equally be replaced with atracheotomy tube, endotracheal tube, nasal pillows, or other means ofmaking a sealed connection between the air delivery means and thesubject's airway.

The microprocessor 16 is programmed to perform the following steps, tobe considered in conjunction with Tables 1 and 2. TABLE 1 FuzzyInference Rules for a first embodiment N Fuzzy Interference Rule FuzzyPhase 1 if size is Zero and rate of Increasing then phase is StartInspiration 2 if size is Small and rate of Increasing then phase isEarly Inspiration Positive change is Slowly 3 if size is Large and raceof Steady then phase is Peak Inspiration Positive change is 4 if size isSmall and rate of Decreasing then phase is Late Inspiration Positivechange is Slowly 5 if size is Zero and rate of Decreasing then phase isStart Expiration change is Fast 6 if size is Small and rate ofDecreasing then phase is Early Expiration Negative change is Slowly 7 ifsize is Large and rate of Steady then phase is Peak Expiration Negativechange is 8 if size is Small and rate of Increasing then phase is LateExpiration Negative change is Slowly 9 if size is Zero and rate ofSteady then phase is Expiratory Pause change is 10 always phase isUnchanged

TABLE 2 Association of phases with fuzzy rules for a first embodiment. NPhase Φ_(N) 1 Start Inspiration 0.0 2 Early Inspiration valuesintermediate between 3 Peak Inspiration 0.0 and 0.5 4 Late Inspiration 5Start Expiration 0.50 6 Early Expiration values intermediate between 7Peak Expiration 0.5 and 1.0 8 Late Expiration 9 Expiratory Pause 10Unchanged Φ

-   1. Set desired target values for the duration of inspiration    TI_(TGT), duration of expiration TE_(TGT), and minute ventilation    V_(TGT). Choose suitable constants P₀ and A_(STD) where P₀ is the    desired end expiratory pressure, and A_(STD) is the desired increase    in pressure above P₀ at end inspiration for a breath of duration    TT_(TGT)=TI_(TGT)+TE_(TGT).-   2. Choose a suitable pressure waveform function Π(Φ), such as that    shown in FIG. 2, such that the desired delivery pressure at phase Φ    will be given by:    P=P ₀ +AΠ(Φ)-    where the amplitude A equals the difference between the end    inspiratory pressure and end expiratory pressure. However, other    waveforms may be suitable for subjects with particular needs.-   3. Initialize the phase Φ in the respiratory cycle to zero, and    initialize the current estimates of actual inspiratory and    expiratory duration TI and TE to TI_(TGT) and TE_(TGT) respectively.-   4. Initialize the rate of change of phase during inspiration ΔΦ_(I)    between sampling intervals of length T to:    ΔΦ+=0.5T/TI _(TGT)-   5. Initialize the rate of change of phase during expiration ΔΦ_(E)    to:    Δ _(E)=0.5T/TE _(TGT)-   6. Measure the instantaneous respiratory airflow f_(RESP).-   7. Calculate the average total breath duration TT=TI+TE-   8. Low pass filter the respiratory airflow with an adjustable time    constant τf, where τf is a fixed small fraction of TT.-   9. Calculate the instantaneous ventilation V, as half the absolute    value of the respiratory airflow:    V=0.5|f _(RESP)|-   10. From the target ventilation V_(TGT) and the measured minute    ventilation V, derive an error term V_(ERR), such that large values    of V_(ERR) indicate inadequate ventilation:    V _(ERR)=∫(V _(TGT) −V)dt-   11. Take V_(BAR) as the result of low pass filtering V with a time    constant τV_(BAR) which is long compared with TT.-   12. Calculate a normalized airflow f_(NORM), where    f _(NORM) =f _(RESP) /V _(BAR).-   13. From f_(NORM), calculate the degree of membership in each of the    fuzzy sets whose membership functions are shown in FIG. 3.-   14. Calculate a normalized rate of change df_(NORM)/dΦ, equal to    df_(NORM)/dt divided by the current estimate of the average    respiratory cycle time TT.-   15. From the normalized rate of change, calculate the degree of    membership in each of the fuzzy sets shown in FIG. 4.-   16. For each row N in Table 1, calculate the degree of membership    g_(N) in the fuzzy set shown in the column labelled Fuzzy Phase, by    applying the fuzzy inference rules shown.-   17. Associate with the result of each of the N rules a phase Φ_(N)    as shown in Table 2, noting that Φ₁₀ is the current phase Φ.-   18. Increase each of the Φ_(N) excepting Φ₁₀ by 0.89 τ/TT, to    compensate for the previous low pass filtering step.-   19. Calculate a new instantaneous phase Φ_(INST) as the angle to the    center of gravity of N unit masses at polar coordinates of radius    g_(N) and angle Φ_(N) revolutions.

20. Calculate the smallest signed difference ΔΦ_(INST) bewteen the phaseestimated in the previous step and the current phase. ΔΦ_(INST) = 1 −(ΔΦ_(INST) − Φ) (Φ_(INST) − Φ > 0.5) ΔΦ_(INST) = Φ_(INST) − Φ + 1(Φ_(INST) − Φ < − 0.5) ΔΦINST = Φ_(INST) − Φ (otherwise)

21. Derive a revised estimate ΔΦ_(REV) equal to a weighted mean of thevalue calculated in the previous step and the average value (ΔΦ_(I) orΔΦ_(E) as appropriate). ΔΦ = (1 − W) ΔΦ_(I) + WΔΦ_(INST) (0 < Φ < 0.5)ΔΦ = (1 − W) ΔΦ_(I) + WΔΦ_(INST) (otherwise)

-    Smaller values of W will cause better tracking of phase if the    subject is breathing regularly, and larger values will cause better    tracking of phase if the subject is breathing irregularly.-   22. Derive a blending fraction B, such that the blending fraction is    unity if the subject's ventilation is well above V_(TGT), zero if    the subject is breathing near or below V_(TGT), and increasing    proportionally from zero to unity as the subject's ventilation    increases through an intermediate range.

23. Calculate ΔΦ_(BLEND) influenced chiefly by ΔΦ calculated in step 21from the subject's respiratory activity if the subject's ventilation iswell above V_(TGT); influenced chiefly by the target respiratoryduration if the subject is breathing near or below V_(TGT); andproportionally between these two amounts if ventilation is in anintermediate range: ΔΦ_(BLEND) = B ΔΦ + 0.5 (1 − B) T/TI_(TGT) (0 < Φ <0.5) ΔΦ_(BLEND) = B ΔΦ + 0.5 (1 − B) T/TE_(TGT) (otherwise)

-   24. Increment Φ by ΔΦ_(BLEND)

25. Update the average rate of change of phase (ΔΦ_(I) or ΔΦ_(E) asappropriate). ΔΦ_(I) = T/τ_(VBAR) (ΔΦ_(BLEND) − ΔΦ_(I)) (0 < Φ < 0.5)ΔΦ_(E) = T/τ_(VBAR) (ΔΦ_(BLEND) − ΔΦ_(E)) (otherwise)

-   26. Recalculate the approximate duration of inspiration TI and    expiration TE:    TI=0.5T/ΔΦ _(I)    TE=0.5T/ΔΦ _(E)

27. Calculate the desired mask pressure modulation amplitude A_(D):A_(D) = A_(STD)/2 (TT < TT_(STD)/2) A_(D) = 2 · A_(STD) (TT > 2 ·TT_(STD)) A_(D) = A_(STD) · TT/TT_(STD) (otherwise)

28. From the error term V_(ERR), calculate an additional mask pressuremodulation amplitude A_(E): A_(E) = K · V_(ERR) (for V_(ERR) > 0) A_(E)= 0 (otherwise)

-    where larger values of K will produce a faster but less stable    control of the degree of assistance, and smaller values of K will    produce slower but more stable control of the degree of assistance.-   29. Set the mask pressure P_(MASK) to:    P _(MASK) =P ₀+(A _(D) +A _(E))Π(Φ)-   30. Wait for a sampling interval T, short compared with the duration    of a respiratory cycle, and then continue at the step of measuring    respiratory airflow.    Measurement of Respiratory Airflow

As follows from above, it is necessary to respiratory airflow, which isa standard procedure to one skilled in the art. In the absence of leak,respiratory airflow can be measured directly with a pneumotachographplaced between the mask and the exhaust. In the presence of a possibleleak, one method disclosed in European Publication No 0 651 971incorporated herein by cross-reference is to calculate the mean flowthrough the leak, and thence calculate the amount of modulation of thepneumotachograph flow signal due to modulation of the flow through theleak induced by changing mask pressure, using the following steps:

-   1. Measure the airflow at the mask f_(MASK) using a pneumotachograph-   2. Measure the pressure at the mask P_(MASK)-   3. Calculate the mean leak as the low pass filtered airflow, with a    time constant long compared with a breath.-   4. Calculate the mean mask pressure as the low pass filtered mask    pressure, with a time constant long compared with a breath.-   5. Calculate the modulation of the flow through the leak as:    δ(leak)=0.5 times the mean leak times the inducing pressure,    where the inducing pressure is P_(MASK)−mean mask pressure.    Thence the instantaneous respiratory airflow can be calculated as:    f _(RESP) =f _(MASK)−mean leak−δ(leak)    A convenient extension as further disclosed in EP 0 651 971    (incorporated herein by cross-reference) is to measure airflow    f_(TURBINE) and pressure P_(TURBINE) at the outlet of the turbine,    and thence calculate P_(MASK) and f_(MASK) by allowing for the    pressure drop down the air delivery hose, and the airflow lost via    the exhaust:-   1. ΔP_(HOS)E=K₁(F_(TURBINE))−K₂(F_(TURBINE))²-   2. _(PMASK)=P_(TURBINE)−ΔP_(HOSE)-   3. F_(EXHAUST)=K3√P_(MASK)-   4. F_(MASK)=F_(TURBINE)−F_(EXHAUST)

Alternative Embodiment

The following embodiment is particularly applicable to subjects withvarying respiratory mechanics, insufficient respiratory drive, abnormalchemoreceptor reflexes, hypoventilation syndromes, or Cheyne Stokesbreathing, or to subjects with abnormalities of the upper or lowerairways, lungs, chest wall, or neuromuscular system.

Many patients with severe lung disease cannot easily be treated using asmooth physiological pressure waveform, because the peak pressurerequired is unacceptably high, or unachievable with for example anose-mask. Such patients may prefer a square pressure waveform, in whichpressure rises explosively fast at the moment of commencement ofinspiratory effort. This may be particularly important in patients withhigh intrinsic PEEP, in which it is not practicable to overcome theintrinsic PEEP by the use of high levels of extrinsic PEEP or CPAP, dueto the risk of hyperinflation. In such subjects, any delay in triggeringis perceived as very distressing, because of the enormous mis-matchbetween expected and observed support. Smooth waveforms exaggerate theperceived delay, because of the time taken for the administered pressureto exceed the intrinsic PEEP. This embodiment permits the use ofwaveforms varying continuously from square (suitable for patients withfor example severe lung or chest wall disease or high intrinsic PEEP) tovery smooth, suitable for patients with normal lungs and chest wall, butabnormal respiratory control, or neuromuscular abnormalities. Thiswaveform is combined either with or without elements of proportionalassist ventilation (corrected for sudden changes in leak), withservo-control of the minute ventilation to equal or exceed a targetventilation. The latter servo-control has an adjustable gain, so thatsubjects with for example Cheyne Stokes breathing can be treated using avery high servo gain to over-ride their own waxing and waning patterns;subjects with various central hypoventilation syndromes can be treatedwith a low servo gain, so that short central apneas are permitted, forexample to cough, clear the throat, talk, or roll over in bed, but onlyif they follow a previous period of high ventilation; and normalsubjects are treated with an intermediate gain.

Restating the Above in Other Words:

-   -   The integral gain of the servo-control of the degree of        assistance is adjustable from very fast (0.3 cmH₂O/L/sec/sec) to        very slow. Patients with Cheyne-Stokes breathing have a very        high ventilatory control loop gain, but a long control loop        delay, leading to hunting. By setting the loop gain even higher,        the patient's controller is stabilized. This prevents the        extreme breathlessness that normally occurs during each cycle of        Cheyne-Stokes breathing, and this is very reassuring to the        patient. It is impossible for them to have a central apnea.        Conversely, subjects with obesity-hypoventilation syndrome have        low or zero loop gain. They will not feel breathless during a        central apnea. However, they have much mucus and need to cough,        and are also often very fidgety, needing to roll about in bed.        This requires that they have central apneas which the machine        does not attempt to treat. By setting the loop gain very low,        the patient is permitted to take a couple of deep breaths and        then have a moderate-length central apnea while coughing,        rolling over, etc, but prolonged sustained apneas or hypopneas        are prevented.    -   Sudden changes in leakage flow are detected and handled using a        fuzzy logic algorithm. The principle of the algorithm is that        the leak filter time constant is reduced dynamically to the        fuzzy extent that the apparent respiratory airflow is a long way        from zero for a long time compared with the patient's expected        respiratory cycle length.    -   Rather than simply triggering between two states (IPAP, EPAP),        the device uses a fuzzy logic algorithm to estimate the position        in the respiratory cycle as a continuous variable. The algorithm        permits the smooth pressure waveform to adjust it's rise time        automatically to the patient's instantaneous respiratory        pattern.    -   The fuzzy phase detection algorithm under normal conditions        closely tracks the patient's breathing. To the extent that there        is a high or suddenly changing leak, or the patient's        ventilation is low, the rate of change of phase (respiratory        rate) smoothly reverts to the specified target respiratory rate.        Longer or deeper hypopneas are permitted to the extent that        ventilation is on average adequate. To the extent that the servo        gain is set high to prevent Cheyne Stokes breathing, shorter and        shallower pauses are permitted.    -   Airflow filtering uses an adaptive filter, which shortens it's        time constant if the subject is breathing rapidly, to give very        fast response times, and lenthens if the subject is breathing        slowly, to help eliminate cardiogenic artifact.    -   The fuzzy changing leak detection algorithm, the fuzzy phase        detection algorithm with its differential handling of brief        expiratory pauses, and handling of changing leak, together with        the smooth waveform severally and cooperatively make the system        relatively immune to the effects of sudden leaks.    -   By suitably setting various parameters, the system can operate        in CPAP, bilevel spontaneous, bilevel timed, proportional assist        ventilation, volume cycled ventilation, and volume cycled        servo-ventilation, and therefore all these modes are subsets of        the present embodiment. However, the present embodiment permits        states of operation that can not be achieved by any of the above        states, and is therefore distinct from them.        Notes        Note 1: in this second embodiment, the names and symbols used        for various quantities may be different to those used in the        first embodiment.        Note 2: The term “swing” is used to refer to the difference        between desired instantaneous pressure at end inspiration and        the desired instantaneous pressure at end expiration.        Note 3: A fuzzy membership function is taken as returning a        value between zero for complete nonmembership and unity for        complete membership. Fuzzy intersection A AND B is the lesser of        A and B, fuzzy union A OR B is the larger of A and B, and fuzzy        negation NOT A is 1−A.        Note 4: root(x) is the square root of x, abs(x) is the absolute        value of x, sign(x) is −1 if x is negative, and +1 otherwise. An        asterisk (*) is used to explicitly indicate multiplication where        this might not be obvious from context.        Apparatus

The apparatus for the second embodiment is shown in FIG. 1 b. The blower110 delivers air under pressure to the mask 111 via the air deliveryhose 112. Exhaled air is exhausted via the exhaust 113 in the mask 111.The pneumotachograph 114 and a differential pressure transducer 115measure the airflow in the nose 112. The flow signal is delivered to themicroprocessor 116. Pressure at any convenient point 117 along the nose112 is measured using a pressure transducer 118. The output from thepressure transducer 118 is delivered to the microcontroller 116 and alsoto a motor servo 119. The microprocessor 116 supplies the motor servo119 with a pressure request signal, which is then compared with thesignal from the pressure transducer 118 to control the blower motor 120.User configurable parameters are loaded into the microprocessor 116 viaa communications port 121, and the computed mask pressure and flow canif desired be output via the communications port 121.

Initialization

The following user adjustable parameters are specified and stored: maxpermissible maximum permissible mask pressure pressure max swing maximumpermissible difference between end inspiratory pressure and endexpiratory pressure. min swing minimum permissible difference betweenend inspiratory pressure and end expiratory pressure. epap endexpiratory pressure min permissible minimum permissible mask pressurepressure target ventilation minute ventilation is sevo-controlled toequal or exceed this quantity target frequency Expected respiratoryrate. If the patient is achieving no respiratory airflow, the pressurewill cycle at this frequency. target duty cycle Expected ratio ofinspiratory time to cycle time. If the patient is achieving norespiratory airflow, the pressure will follow this duty cycle. linearresistance resistive unloading = linear resistance * f + andquadresistance quad_resistance * f² sign(f), where f is the respiratoryairflow. where sign(x) = −1 for x < 0, +1 otherwise elastance Unload atleast this much elastance servo gain gain for servo-control of minuteventilation to at least exceed target ventilation. waveform timeconstant Elastic unloading waveform time constant as a fraction ofinspiratory duration. (0.0 = square wave) hose resistance ΔP frompressure sensing port to inside mask = hose resistance times the squareof the flow in the intervening tubing. diffuser conductance Flow throughthe mask exhaust port = diffuser conductance * root mask pressureAt initialization, the following are calculated from the aboveuser-specified settings:The expected duration of a respiratory cycle, of an inspiration, and ofan expiration are set respectively to:STDT _(TOT)=60/target respiratory rateSTDT _(I) =STDT _(TOT)*target duty cycleSTDT _(E) =STDT _(TOT) −STDT _(I)The standard rates of change of phase (revolutions per sec) duringinspiration and expiration are set respectively to:STDdφ _(I)=0.5/STDT _(I)STDdφ _(E)=0.5/STDT _(E)The instantaneous elastic support at any phase φ in the respiratorycycle is given by:PEL(φ)=swing*Π(φ)

where swing is the pressure at end inspiration minus the pressure at endexpiration, π(Φ) = e⁻²τΦ during inspiration, e⁻⁴t(Φ − 0.5) duringexpirationand τ is the user-selectable waveform time constant.If τ=0, then Π(φ) is a square wave. The maximum implemented value forτ=0.3, producing a waveform approximately as shown in FIG. 5.The mean value of Π(φ) is calculated as follows:$\Pi_{BAR} = {0.5{\int_{0}^{.05`}{{\Pi(\phi)}\quad{\mathbb{d}\phi}}}}$Operations Performed Every 20 Milliseconds

The following is an overview of routine processing done at 50 Hz:

-   -   measure flow at flow sensor and pressure at pressure sensing        port    -   calculate mask pressure and flow from sensor pressure and flow    -   calculate conductance of mask leak    -   calculate instantaneous airflow through leak    -   calculate respiratory airflow and low pass filtered respiratory        airflow    -   calculate mask on-off status and lead-in    -   calculate instantaneous and recent peak jamming    -   calculate time constant for leak conductance calculations    -   calculate phase in respiratory cycle    -   update mean rates of change of phase for inspiration and        expiration, lengths of inspiratory and expiratory times, and        respiratory rate    -   add hose pressure loss to EPAP pressure    -   add resistive unloading    -   calculate instantaneous elastic assistance required to        servo-control ventilation    -   estimate instantaneous elastic recoil pressure using various        assumptions    -   weight and combine estimates    -   add servo pressure to yield desired sensor pressure    -   servo-control motor speed to achieve desired sensor pressure        The details of each step will now be explained.        Measurement of Flow and Pressure        Flow is measured at the outlet of the blower using a        pneumotachograph and differential pressure transducer. Pressure        is measured at any convenient point between the blower outlet        and the mask. A humidifier and/or anti-bacterial filter may be        inserted between the pressure sensing port and the blower. Flow        and pressure are digitized at 50 Hz using an A/D converter.        Calculation of Mask Flow and Pressure        The pressure loss from pressure measuring point to mask is        calculated from the flow at the blower and the (quadratic)        resistance from measuring point to mask.        Hose pressure loss=sign(flow)*hose resistance*flow²        where sign(x)=−1 for x<0, +1 otherwise. The mask pressure is        then calculated by subtracting the hose pressure loss from the        measured sensor pressure:        Mask pressure=sensor pressure−hose pressure loss        The flow through the mask exhaust diffuser is calculated from        the known parabolic resistance of the diffuser holes, and the        square root of the mask pressure:        diffuser flow=exhaust resistance*sign(mask        pressure)*root(abs(mask pressure))        Finally, the mask flow is calculated:        mask flow=sensor flow−diffuser flow        The foregoing describes calculation of mask pressure and flow in        the various treatment modes. In diagnostic mode, the patient is        wearing only nasal cannulae, not a mask. The cannula is plugged        into the pressure sensing port. The nasal airflow is calculated        from the pressure, after a linearization step, and the mask        pressure is set to zero by definition.        Conductance of Leak        The conductance of the leak is calculated as follows:        root mask pressure=sign(P_(MASK))√{square root over        (abs(P_(MASK)))}        LP mask airflow=low pass filtered mask airflow        LP root mask pressure=low pass filtered root mask pressure        conductance of leak=LP mask airflow/LP root mask pressure        The time constant for the two low pass filtering steps is        initialized to 10 seconds and adjusted dynamically thereafter        (see below).        Instantaneous Flow Through Leak        The instantaneous flow through the leak is calculated from the        instantaneous mask pressure and the conductance of the leak:        instantaneous leak=conductance of leak*root mask pressure        Respiratory Airflow        The respiratory airflow is the difference between the flow at        the mask and the instantaneous leak:        respiratory airflow=maskflow−instantaneous leak        Low Pass Filtered Respiratory Airflow        Low pass filter the respiratory airflow to remove cardiogenic        airflow and other noise. The time constant is dynamically        adjusted to be 1/40 of the current estimated length of the        respiratory cycle T_(TOT) (initialized to STD_T_(TOT) and        updated below). This means that at high respiratory rates, there        is only a short phase delay introduced by the filter, but at low        respiratory rates, there is good rejection of cardiogenic        airflow.        Mask On/Off Status        The mask is assumed to initially be off. An off-on transition is        taken as occurring when the respiratory airflow first goes above        0.2 L/sec, and an on-off transition is taken as occurring if the        mask pressure is less than 2 cmH₂O for more than 1.5 seconds.        Lead-In        Lead-in is a quantity that runs from zero if the mask is off, or        has just been donned, to 1.0 if the mask has been on for 20        seconds or more, as shown in FIG. 6.        Calculation of Instantaneous Jamming Index, J        J is the fuzzy extent to which the impedance of the leak has        suddenly changed. It is calculated as the fuzzy extent to which        the absolute magnitude of the respiratory airflow is large for        longer than expected.        The fuzzy extent A_(I) to which the airflow has been positive        for longer than expected is calculated from the time t_(ZI)        since the last positive-going zero crossing of the calculated        respiratory airflow signal, and the expected duration STD T_(I)        of a normal inspiration for the particular subject, using the        fuzzy membership function shown in FIG. 7.        The fuzzy extent B_(I) to which the airflow is large and        positive is calculated from the instantaneous respiratory        airflow using the fuzzy membership function shown in FIG. 8.        The fuzzy extent I_(I) to which the leak has suddenly increased        is calculated by calculating the fuzzy intersection (lesser) of        A_(I) and B_(I).        Precisely symmetrical calculations are performed for expiration,        deriving I_(E).as the fuzzy extent to which the leak has        suddenly decreased. A_(E) is calculated from T_(ZE) and T_(E),        B_(E) is calculated from minus f_(RESP), and I_(E) is the fuzzy        intersection of A_(E) and B_(E). The instantaneous jamming index        J is calculated as the fuzzy union (larger) of indices I_(I) and        I_(E).        Recent Peak Jamming        If the instantaneous jamming index is larger than the current        value of the recent peak jamming index, then the recent peak        jamming index is set to equal the instantaneous jamming index.        Otherwise, the recent peak jamming index is set to equal the        instantaneous jamming index low pass filtered with a time        constant of 10 seconds. An electrical analogy of the calculation        is shown in FIG. 9.        Time Constant for Leak Conductance Calculations        If the conductance of the leak suddenly changes, then the        calculated conductance will initially be incorrect, and will        gradually approach the correct value at a rate which will be        slow if the time constant of the low pass filters is long, and        fast if the time constant is short. Conversely, if the impedance        of the leak is steady, the longer the time constant the more        accurate the calculation of the instantaneous leak. Therefore,        it is desirable to lengthen the time constant to the extent that        the leak is steady, reduce the time constant to the extent that        the leak has suddenly changed, and to use intermediately longer        or shorter time constants if it is intermediately the case that        the leak is steady.        If there is a large and sudden increase in the conductance of        the leak, then the calculated respiratory airflow will be        incorrect. In particular, during apparent inspiration, the        calculated respiratory airflow will be large positive for a time        that is large compared with the expected duration of a normal        inspiration. Conversely, if there is a sudden decrease in        conductance of the leak, then during apparent expiration the        calculated respiratory airflow will be large negative for a time        that is large compared with the duration of normal expiration.        Therefore, the time constant for the calculation of the        conductance of the leak is adjusted depending on J_(PEAK), which        is a measure of the fuzzy extent that the leak has recently        suddenly changed, as shown in FIG. 10.        In operation, to the extent that there has recently been a        sudden and large change in the leak, J_(PEAK) will be large, and        the time constant for the calculation of the conductance of the        leak will be small, allowing rapid convergence on the new value        of the leakage conductance. Conversely, if the leak is steady        for a long time, J_(PEAK) will be small, and the time constant        for calculation of the leakage conductance will be large,        enabling accurate calculation of the instantaneous respiratory        airflow. In the spectrum of intermediate situations, where the        calculated instantaneous respiratory airflow is larger and for        longer periods, J_(PEAK) will be progressively larger, and the        time constant for the calculation of the leak will progressively        reduce. For example, at a moment in time where it is uncertain        whether the leak is in fact constant, and the subject has merely        commenced a large sigh, or whether in fact there has been a        sudden increase in the leak, the index will be of an        intermediate value, and the time constant for calculation of the        impedance of the leak will also be of an intermediate value. The        advantage is that some corrective action will occur very early,        but without momentary total loss of knowledge of the impedance        of the leak.        Instantaneous Phase in Respiratory Cycle        The current phase φ runs from 0 for start of inspiration to 0.5        for start of expiration to 1.0 for end expiration=start of next        inspiration. Nine separate features (peaks, zero crossings,        plateaux, and some intermediate points) are identified on the        waveform, as shown in FIG. 11.        Calculation of Normalized Respiratory Airflow        The filtered respiratory airflow is normalized with respect to        the user specified target ventilation as follows:        standard airflow=target ventilation/7.5 L/min        f′=filtered respiratory airflow/standard airflow        Next, the fuzzy membership in fuzzy sets large negative, small        negative, zero, small positive, and large positive, describing        the instantaneous airflow is calculated using the membership        functions shown in FIG. 12. For example, if the normalized        airflow is 0.25, then the airflow is large negative to extent        0.0, small negative to extent 0.0, zero to extent 0.5, small        positive to extent 0.5, large positive to extent 0.00.        Calculation of Normalized Rate of Change of Airflow        The rate of change of filtered respiratory airflow is calculated        and normalized to a target ventilation of 7.5 L/min at 15        breaths/min as follows:        standard df/dt=standard airflow*target frequency/15        calculate d(filtered airflow)/dt        low pass filter with a time constant of 8/50 seconds        normalize by dividing by standard df/dt        Now evaluate the membership of normalized df/dt in the fuzzy        sets falling, steady, and rising, whose membership functions are        shown in FIG. 13.        Calculation of Ventilation, Normalized Ventilation, and Hypopnea        ventilation=abs(respiratory airflow), low pass filtered with a        time constant of STDT _(TOT).        normalized ventilation=ventilation/standard airflow        Hypopnea is the fuzzy extent to which the normalized ventilation        is zero. The membership function for hypopnea is shown in FIG.        14.        Calculation of Recent Ventilation, Normalized Recent        Ventilation, and Hyperpnea        Recent ventilation is also a low pass filtered abs(respiratory        airflow), but filtered with an adjustable time constant,        calculated from servo gain (specified by the user) as shown in        FIG. 15. For example, if the servo gain is set to the maximum        value of 0.3, the time constant is zero, and recent ventilation        equals instantaneous abs(respiratory airflow). Conversely, if        servo gain is zero, the time constant is twice STD T_(TOT), the        expected length of a typical breath.        Target absolute airflow=2*target ventilation        normalized recent ventilation=recent ventilation/target absolute        airflow        Hyperpnea is the fuzzy extent to which the recent ventilation is        large. The membership function for hyperpnea is shown in FIG.        16.        Big Leak        The fuzzy extent to which there is a big leak is calculated from        the membership function shown in FIG. 17.        Additional Fuzzy Sets Concerned with Fuzzy “Triggering”        Membership in fuzzy sets switch negative and switch positive are        calculated from the normalized respiratory airflow using the        membership functions shown in FIG. 18, and membership in fuzzy        sets insp_phase and exp_phase are calculated from the current        phase f using the membership functions shown in FIG. 19.        Fuzzy Inference Rules for Phase        Procedure W(y) calculates the area of an isosceles triangle of        unit height and unit base, truncated at height y as shown in        FIG. 20. In the calculations that follow, recall that fuzzy        intersection a AND b is the smaller of a and b, fuzzy union a OR        b is the larger of a and b, and fuzzy negation NOT a is 1−a.        The first fuzzy rule indicates that lacking any other        information the phase is to increase at a standard rate. This        rule is unconditionally true, and has a very heavy weighting,        especially if there is a large leak, or there has recently been        a sudden change in the leak, or there is a hypopnea.        W _(STANDARD)=8+16*J _(PEAK)+16*hyopopnea+16*big leak        The next batch of fuzzy rules correspond to the detection of        various features of a typical flow-vs-time curve. These rules        all have unit weighting, and are conditional upon the fuzzy        membership in the indicated sets:        W _(EARLY INSP) =W(rise and small positive)        WPEAK INSP=W(large positive AND steady AND NOT recent peak        jamming)        W _(LATE INSP) =W(fall AND small positive)        W _(EARLY EXP) =W(fall AND small negative)        W _(PEAK EXP) =W(large negative AND steady)        W _(LATE EXP) =W(rise AND small negative)        The next rule indicates that there is a legitimate expiratory        pause (as opposed to an apnea) if there has been a recent        hyperpnea and the leak has not recently changed:        W _(PAUSE)=(hyperpnea AND NOT J _(PEAK))*W(steady AND zero)        Recalling that the time constant for hyperpnea gets shorter as        servo gain increases, the permitted length of expiratory pause        gets shorter and shorter as the servo gain increases, and        becomes zero at maximum servo gain. The rationale for this is        that (i) high servo gain plus long pauses in breathing will        result in “hunting” of the servo-controller, and (ii) in general        high servo gain is used if the subject's chemoreceptor responses        are very brisk, and suppression of long apneas or hypopneas will        help prevent the subject's own internal servo-control from        hunting, thereby helping prevent Cheyne-Stokes breathing.        Finally, there are two phase-switching rules. During regular        quiet breathing at roughly the expected rate, these rules should        not strongly activate, but they are there to handle irregular        breathing or breathing at unusual rates. They have very heavy        weightings.        W _(TRIG INSP)=32W(expiratory phase AND switch positive)        W _(TRIG EXP)=32W(inspiratory phase AND switch negative)        Defuzzification

For each of the ten fuzzy rules above, we attach phase angles f_(N), asshown in Table ZZZ. Note that φ are in revolutions, not radians. We nowplace the ten masses W(N) calculated above at the appropriate phaseangles φ_(N) around the unit circle, and take the centroid. Rule N Φ_(N)STANDARD 1 current Φ TRIG INSP 2 0.00 EARLY INSP 3 0.10 PEAK INSP 4 0.30LATE INSP 5 0.50 TRIG EXP 6 0.5 + 0.05 k EARLY EXP 7 0.5 + 0.10 k PEAKEXP 8 0.5 + 0.20 k LATE EXP 9 0.5 + 0.4 k EXP PAUSE 10 0.5 + 0.5 kwhere k=STD T_(I)/STD T_(E).Note that if the user has entered very short duty cycle, k will besmall. For example a normal duty cycle is 40%, giving k= 40/60=0.67.Thus the expiratory peak will be associated with a phase angle of0.5+0.2*0.67=0.63, corresponding 26% of the way into expiratory time,and the expiratory pause would start at 0.5+0.5*0.67=0.83, correspondingto 67% of the way into expiratory time. Conversely, if the duty cycle isset to 20% in a patient with severe obstructive lung disease, features 6through 10 will be skewed or compressed into early expiration,generating an appropriately longer expiratory pause.The new estimate of the phase is the centroid, in polar coordinates, ofthe above ten rules:${centroid} = {{arc}\quad{\tan( \frac{\sum{W_{N}\sin\quad\phi_{N}}}{\sum{W_{N}\cos\quad\phi_{N}}} )}}$The change in phase dφ from the current phase φ to the centroid iscalculated in polar coordinates. Thus if the centroid is 0.01 and thecurrent phase is 0.99, the change in phase is dφ=0.02. Conversely, ifthe centroid is 0.99 and the current phase is 0.01, then dφ=−0.02. Thenew phase is then set to the centroid:φ=centroidThis concludes the calculation of the instantaneous phase in therespiratory cycle φ.Estimated Mean Duration of Inspiration, Expiration, Cycle Time, andRespiratory RateIf the current phase is inspiratory (φ<0.5) the estimated duration ofinspiration T_(I) is updated:LP(dφ _(I))=low pass filtered dφ with a time constant of 4*STDT _(TOT)Clip LP(dφ_(I)) to the range (0.5/STDT_(I))/2 to 4(0.5/STDT_(I))T _(I)=0.5/clipped LP(dφI)Conversely, if the current phase is expiratory, (φ>=0.5) the estimatedduration of expiration T_(E) is updated:LP(dφ _(E))=low pass filtered dφ with a time constant of 4*STDT _(TOT)Clip LP(dφE) to the range (0.5/STDT_(E))/2 to 4(0.5/STDT_(E))T _(E)=0.5/clipped LP(dφ _(E))The purpose of the clipping is firstly to prevent division by zero, andalso so that the calculated T_(I) and T_(E) are never more than a factorof 4 shorter or a factor of 2 longer than expected.Finally, the observed mean duration of a breath T_(TOT) and respiratoryrate RR are:T _(TOT) =T _(I) +T _(E)RR=60/T _(TOT)Resistive UnloadingThe resistive unloading is the pressure drop across the patient's upperand lower airways, calculated from the respiratory airflow andresistance values stored in SRAMf=respiratory airflow truncated to +/−2 L/secresistive unloading=airway resistance*f+upper airway resistance*f²*sign(f)Instantaneous Elastic AssistanceThe purpose of the instantaneous elastic assistance is to provide apressure which balances some or all of the elastic deflating pressuresupplied by the springiness of the lungs and chest wall (instantaneouselastic pressure), plus an additional component required toservo-control the minute ventilation to at least exceed on average apre-set target ventilation. In addition, a minimum swing, alwayspresent, is added to the total. The user-specified parameter elastanceis preset to say 50-75% of the known or estimated elastance of thepatient's lung and chest wall. The various components are calculated asfollows:Instantaneous Assistance Based on Minimum Pressure Swing Set byPhysician:instantaneous minimum assistance=minimum swing*Π(φ)Elastic Assistance Required to Servo-Control Ventilation to Equal orExceed TargetThe quantity servo swing is the additional pressure modulation amplituderequired to servo-control the minute ventilation to at least equal onaverage a pre-set target ventilation.Minute ventilation is defined as the total number of litres inspired orexpired per minute. However, we can't wait for a whole minute, or evenseveral seconds, to calculate it, because we wish to be able to preventapneas or hypopneas lasting even a few seconds, and a PI controllerbased on an average ventilation over a few seconds would be eithersluggish or unstable.The quantity actually servo-controlled is half the absolute value of theinstantaneous respiratory airflow. A simple clipped integral controllerwith no damping works very satisfactorily. The controller gain andmaximum output ramp up over the first few seconds after putting the maskon.If we have had a sudden increase in mouth leak, airflow will be nonzerofor a long time. A side effect is that the ventilation will be falselymeasured as well above target, and the amount of servo assistance willbe falsely reduced to zero. To prevent this, to the extent that thefuzzy recent peak jamming index is large, we hold the degree of servoassistance at its recent average value, prior to the jamming.The algorithm for calculating servo swing is as follows:error=target ventilation−abs(respiratory airflow)/2servo swing=Serror*servo gain*sample intervalclip servo swing to range 0 to 20 cmH₂O*lead-inset recent servo swing=servo swing low pass filtered with a timeconstant of 25 sec.clip servo swing to be at most J_(PEAK)*recent servo swingThe instantaneous servo assistance is calculated by multiplying servoswing by the previously calculated pressure waveform template:instantaneous servo assistance=servo swing*Π(φ)Estimating Instantaneous Elastic PressureThe instantaneous pressure required to unload the elastic work ofinspiring against the user-specified elastance is the specifiedelastance times the instantaneous inspired volume. Unfortunately,calculating instantaneous inspired volume simply by integratingrespiratory airflow with respect to time does not work in practice forthree reasons: firstly leaks cause explosive run-away of theintegration. Secondly, the integrator is reset at the start of eachinspiration, and this point is difficult to detect reliably. Thirdly,and crucially, if the patient is making no efforts, nothing will happen.Therefore, four separate estimates are made, and a weighted averagetaken.Estimate 1: Exact instantaneous elastic recoil calculated frominstantaneous tidal volume, with a correction for sudden change in leakThe first estimate is the instantaneous elastic recoil of a specifiedelastance at the estimated instantaneous inspired volume, calculated bymultiplying the specified elastance by the integral of a weightedrespiratory airflow with respect to time, reset to zero if therespiratory phase is expiratory. The respiratory airflow is weighted bythe fuzzy negation of the recent peak jamming index J_(PEAK), to partlyameliorate an explosive run-away of the integral during brief periods ofsudden increase in leak, before the leak detector has had time to adaptto the changing leak. In the case where the leak is very steady,J_(PEAK) will be zero, the weighting will be unity, and the inspiredvolume will be calculated normally and correctly. In the case where theleak increases suddenly, J_(PEAK) will rapidly increase, the weightingwill decrease, and although typically the calculated inspired volumewill be incorrect, the over-estimation of inspired volume will beameliorated. Calculations are as follows:Instantaneous volume=integral of respiratory airflow*(1−J _(PEAK))dtif phase is expiratory (0.5<φ<1.0 revolutions) reset integral to zeroestimate 1=instantaneous volume*elastanceEstimate 2: based on assumption that the tidal volume equals the targettidal volumeThe quantity standard swing is the additional pressure modulationamplitude that would unload the specified elastance for a breath of apreset target tidal volume.target tidal volume=target ventilation/target frequencystandard swing=elastance*target tidal volumeestimate 2=standard swing*Π(φ)Estimate 3: based on assumption that the tidal volume equals the targettidal volume divided by the observed mean respiratory rate RR calculatedpreviously.Estimate 3=elastance*target ventilation/RR*Π(φ)Estimate 4: based on assumption that this breath is much like recentbreathsThe instantaneous assistance based on the assumption that the elasticwork for this breath is similar to that for recent breaths is calculatedas follows:LP elastic assistance=instantaneous elastic assistance low pass filteredwith a time constant of 2STDT _(TOT)estimate 4=LP elastic assistance*Π(φ)/P _(BAR)

The above algorithm works correctly even if Π(φ) is dynamically changedon-the-fly by the user, from square to a smooth or vice versa. Forexample, if an 8 cmH₂O square wave (Π_(BAR)=1) adequately assists thepatient, then a sawtooth wave (Π_(BAR)=0.5) will require 16 cmH₂O swingto produce the same average assistance.

Best Estimate of Instantaneous Elastic Recoil Pressure

Next, calculate the pressure required to unload a best estimate of theactual elastic recoil pressure based on a weighted average of the above.If Π(φ) is set to the smoothest setting, the estimate is based equallyon all the above estimates of instantaneous elastic recoil. If Π(φ) is asquare wave, the estimate is based on all the above estimates except forestimate 1, because a square wave is maximal at φ=0, whereas estimate 1is zero at φ=0. Intermediate waveforms are handled intermediately.Quantity smoothness runs from zero for a square wave to 1 for a waveformtime constant of 0.3 or above.smoothness=waveform time constant/0.3instantaneous recoil=(smoothness*estimate 1+estimate 2+estimate3+estimate 4)/(smoothness+3)Now add the estimates based on minimum and servo swing, truncate so asnot to exceed a maximum swing set by the user. Reduce (lead ingradually) if the mask has only just been put on.I=instantaneous minimum assistance+instantaneous servoassistance+instantaneous recoilTruncate I to be less than preset maximum permissible swinginstantaneous elastic assistance=I*lead-inThis completes the calculation of instantaneous elastic assistance.Desired Pressure at Sensordesired sensor pressure=epap+hose pressure loss+resistiveunloading+instantaneous elastic assistanceServo Control of Motor Speed

In the final step, the measured pressure at the sensor isservo-controlled to equal the desired sensor pressure, using for examplea clipped pseudodifferential controller to adjust the motor current.Reference can be made to FIG. 1 in this regard.

Device Performance

FIGS. 21-27 each show an actual 60 second recording displaying an aspectof the second embodiment. All recordings are from a normal subjecttrained to perform the required manoeuvres. Calculated respiratoryairflow, mask pressure, and respiratory phase are calculated using thealgorithms disclosed above, output via a serial port, and plotteddigitally.

In FIGS. 21-26 respiratory airflow is shown as the darker tracing, thevertical scale for flow being ±L/sec, inspiration upwards. The verticalscale for the pressure (light trace) is 0.2 cmH₂O.

FIG. 21 is recorded with the servo gain set to 0.1 cmH₂O/L/sec/sec,which is suitable for subjects with normal chemoflexes. The subject isbreathing well above the minimum ventilation, and a particularly deepbreath (sigh) is taken at point (a). As is usual, respiratory effortceases following the sigh, at point (c). The device correctly permits ashort central apnea (b), as indicated by the device remaining at the endexpiratory pressure during the period marked (b). Conversely FIG. 22shows that if there is no preceding deep breath, when efforts cease at(a), the pressure correctly continues to cycle, thus preventing anyhypoxia. FIG. 23 is recorded with servo gain set high, as would beappropriate for a subject with abnormally high chemoreflexes such as istypically the case with Cheyne-Stokes breathing. Now when effort ceasesat arrow (a), pressure continues to cycle and a central apnea is nolonger permitted, despite preceding deep breathing. This is advantageousfor preventing the next cycle of Cheyne-Stokes breathing.

The above correct behaviour is also exhibited by a time mode device, butis very different to that of a spontaneous mode bilevel device, orequally of proportional assist ventilation, both of which would fail tocycle after all central apneas, regardless of appropriateness.

FIG. 24 shows automatically increasing end-inspiratory pressure as thesubject makes voluntarily deeper inspiratory efforts. The desirablebehaviour is in common with PAV, but is different to that of a simplebilevel device, which would maintain a constant level of support despitean increased patient requirement, or to a volume cycled device, whichwould actually decrease support at a time of increasing need.

FIG. 25 is recorded with a somewhat more square waveform selected. Thisfigure shows automatically increasing pressure support when the subjectvoluntarily attempts to resist by stiffening the chest wall at point(a). This desirable behaviour is common with PAV and volume cycleddevices, with the expectation that PAV cannot selectively deliver asquarer waveform. It is distinct from a simple bilevel device whichwould not augment the level of support with increasing need.

FIG. 26 shows that with sudden onset of a severe 1.4 L/sec leak at (a),the flow signal returns to baseline (b) within the span of a singlebreath, and pressure continues to cycle correctly throughout. Althoughtimed mode devices can also continue to cycle correctly in the face ofsudden changing leak, the are unable to follow the subject's respiratoryrate when required (as shown in FIG. 27). Other known bilevel devicesand PAV mis-trigger for longer or shorter periods following onset of asudden sever leak, and PAV can deliver greatly excessive pressures underthese conditions.

FIG. 27 shows an actual 60 second tracing showing respiratory airflow(heavy trace ±1 L/sec full scale) and respiratory phase as a continuousvariable (light trace, 0 to 1 revolution), with high respiratory rate inthe left half of the trace and low respiratory rate in the right half ofthe trace. This trace demonstrates that the invention can determinephase as a continuous variable.

Advantageous Aspects of Embodiments of the Invention

Use of Phase as a Continuous Variable.

In the prior art, phase is taken as a categorical variable, with twovalues: inspiration and expiration. Errors in the detection of start ofinspiration and start of expiration produce categorical errors indelivered pressure. Conversely, here, phase is treated as a continuousvariable having values between zero and unity. Thus categorical errorsin measurement of phase are avoided.

Adjustable Filter Frequency and Allowance for Phase Delay

By using a short time constant when the subject is breathing rapidly,and a long time constant when the subject is breathing slowly, thefilter introduces a fixed phase delay which is always a small fractionof a respiratory cycle. Thus unnecessary phase delays can be avoided,but cardiogenic artifact can be rejected in subjects who are breathingslowly. Furthermore, because phase is treated as a continuous variable,it is possible to largely compensate for the delay in the low passfilter.

Within-Breath Pressure Regulation as a Continuous Function ofRespiratory Phase.

With all prior art there is an intrusive discontinuous change inpressure, either at the start of inspiration or at the start ofexpiration. Here, the pressure change is continuous, and therefore morecomfortable.

With proportional assist ventilation, the instantaneous pressure is afunction of instantaneous volume into the breath. This means that asudden large leak can cause explosive pressure run-away. Here, whereinstantaneous pressure is a function of instantaneous phase rather thantidal volume, this is avoided.

Between-Breath Pressure-Regulation as a Function of Average InspiratoryDuration.

Average inspiratory duration is easier to calculate in the presence ofleak than is tidal volume. By taking advantage of a correlation betweenaverage inspiratory duration and average tidal volume, it is possible toadjust the amplitude of modulation to suit the average tidal volume.

Provision of a Pressure Component for Unloading Turbulent Upper AirwayResistance, and Avoiding Cardiogenic Pressure Instabilities.

Although Younes describes the use of a component of pressureproportional to the square of respiratory airflow to unload theresistance of external apparatus, the resistance of the externalapparatus in embodiments of the present invention is typicallynegligible. Conversely, embodiments of the present invention describestwo uses for such a component proportional to the square of respiratoryairflow that were not anticipated by Younes. Firstly, sleeping subjects,and subjects with a blocked nose, have a large resistance proportionalto the square of airflow, and a pressure component proportional to thesquare of airflow can be used to unload the anatomical upper airwayresistance. Secondly, small nonrespiratory airflow components due toheartbeat or other artifact, when squared, produces negligible pressuremodulation, so that the use of such a component yields relative immunityto such nonrespiratory airflow.

Smooth Transition Between Spontaneous and Controlled Breathing

There is a smooth, seamless gradation from flexibly tracking thesubject's respiratory pattern during spontaneous breathing well abovethe target ventilation, to fully controlling the duration, depth, andphase of breathing if the subject is making no efforts, via atransitional period in which the subject can make progressively smallerchanges to the timing and depth of breathing. A smooth transition avoidscategorization errors when ventilation is near but not at the desiredthreshold. The advantage is that the transition from spontaneous tocontrolled ventilation occurs unobtrusively to the subject. This can beespecially important in a subject attempting to go to sleep. A similarsmooth transition can occur in the reverse direction, as a subjectawakens and resumes spontaneous respiratory efforts.

1. A method for calculating the instantaneous phase in the respiratorycycle comprising at least the step of determining that if theinstantaneous airflow is small and increasing fast, then it is close tostart of inspiration, if the instantaneous airflow is large and steady,then it is close to mid-inspiration, if the instantaneous airflow issmall and decreasing fast, then it is close to mid-expiration, if theinstantaneous airflow is zero and steady, then it is during anend-expiratory pause, and airflow conditions intermediate between theabove are associated with correspondingly intermediate phases.
 2. Amethod for determining the instantaneous phase in the respiratory cycleas a continuous variable from 0 to 1 revolution, the method comprisingthe steps of: selecting at least two identifiable features F_(N) of aprototype flow-vs-time waveform f(t) similar to an expected respiratoryflow-vs-time waveform, and for each said feature: determining byinspection the phase φ_(N) in the respiratory cycle for said feature,assigning a weight W_(N) to said phase, defining a “magnitude” fuzzy setM_(N) whose membership function is a function of respiratory airflow,and a “rate of change” fuzzy set C_(N), whose membership function is afunction of the time derivative of respiratory airflow, chosen such thatthe fuzzy intersection M_(N) AND C_(N) will be larger for points on thegeneralized prototype respiratory waveform whose phase is closer to thesaid feature F_(N) than for points closer to all other selectedfeatures, setting the fuzzy inference rule R_(N) for the selectedfeature F_(N) to be: If flow is M_(N) and rate of change of flow isC_(N) then phase=φ_(N), with weight W_(N). measuring leak-correctedrespiratory airflow, for each feature F_(N) calculating fuzzy membershipin fuzzy sets M_(N) and C_(N), for each feature F_(N) applying fuzzyinference rule R_(N) to determine the fuzzy extent Y_(N)=M_(N) AND C_(N)to which the phase is φ_(N), and applying a defuzzification procedureusing Y_(N) at phases φ_(N) and weights W_(N) to determine theinstantaneous phase φ.
 3. A method as claimed in claim 2, whereby theidentifiable features include zero crossings, peaks, inflection pointsor plateaus of the prototype flow-vs-time waveform.
 4. A method asclaimed in claim 2, whereby said weights can be unity, or chosen toreflect the anticipated reliability of deduction of the particularfeature.
 5. A method as claimed in claim 2, in which the step ofcalculating respiratory airflow includes a low pass filtering step toreduce non-respiratory noise, in which the time constant of the low passfilter is an increasing function of an estimate of the length of therespiratory cycle.
 6. A method as claimed in claim 2, in which thedefuzzification step includes a correction for any phase delayintroduced in the step of low pass filtering respiratory airflow.
 7. Amethod for measuring the average respiratory rate, comprising the stepsof: determining leak-corrected respiratory airflow, from the respiratoryairflow, calculating the instantaneous phase φ in the respiratory cycleas a continuous variable from 0 to 1 revolution, calculating theinstantaneous rate of change of phase dφ/dt, and calculating the averagerespiratory rate by low pass filtering said instantaneous rate of changeof phase dφ/dt.
 8. A method as claimed in claim 7, whereby theinstantaneous phase is determined by the method of claim 1 or claim 2.9. A method for providing ventilatory assistance in a spontaneouslybreathing subject, comprising the steps, performed at repeated samplingintervals, of: ascribing a desired waveform template function Π(φ), withdomain 0 to 1 revolution and range 0 to 1, calculating the instantaneousphase φ in the respiratory cycle as a continuous variable from 0 to 1revolution, selecting a desired pressure modulation amplitude A,calculating a desired instantaneous delivery pressure as an endexpiratory pressure plus the desired pressure modulation amplitude Amultiplied by the value of the waveform template function Π(φ) at thesaid calculated phase φ, and setting delivered pressure to subject tothe desired delivery pressure.
 10. A method as claimed in claim 9,whereby step of selecting a desired pressure modulation amplitude is afixed amplitude.
 11. A method as claimed in claim 9, whereby the step ofselecting a desired pressure modulation amplitude in which saidamplitude is equal to an elastance multiplied by an estimate of thesubject's tidal volume.
 12. A method for providing ventilatoryassistance in a spontaneously breathing subject as described above, inwhich the step of selecting a desired pressure modulation amplitudecomprises the substeps of: specifying a typical respiratory rate givinga typical cycle time, specifying a preset pressure modulation amplitudeto apply at said typical respiratory rate, calculating the observedrespiratory rate giving an observed cycle time, and calculating thedesired amplitude of pressure modulation as said preset pressuremodulation amplitude multiplied by said observed cycle time divided bythe said specified cycle time.
 13. A method for providing ventilatoryassistance in a spontaneously breathing subject, including at least thestep of determining the extent that the subject is adequatelyventilated, to said extent the phase in the respiratory cycle isdetermined from the subject's respiratory airflow, but to the extentthat the subject's ventilation is inadequate, the phase in therespiratory cycle is assumed to increase at a pre-set rate, and settingmask pressure as a function of said phase.
 14. A method for providingventilatory assistance in a spontaneously breathing subject, comprisingthe steps of: measuring respiratory airflow, determining the extent towhich the instantaneous phase in the respiratory cycle can be determinedfrom said airflow, to said extent determining said phase from saidairflow but to the extent that the phase in the respiratory cycle cannotbe accurately determined, the phase is assumed to increase at a presetrate, and delivering pressure as a function of said phase.
 15. A methodfor calculating the instantaneous inspired volume of a subject, operablesubstantially without run-away under conditions of suddenly changingleak, the method comprising the steps of: determining respiratoryairflow approximately corrected for leak, calculating an index J varyingfrom 0 to 1 equal to the fuzzy extent to which said correctedrespiratory airflow is large positive for longer than expected, or largenegative for longer than expected, identifying the start of inspiration,and calculating the instantaneous inspired volume as the integral ofsaid corrected respiratory airflow multiplied by the fuzzy negation ofsaid index J with respect to time, from start of inspiration.
 16. Amethod for providing ventilatory assistance in a spontaneously breathingsubject, the method comprising the steps, performed at repeated samplingintervals, of: determining respiratory airflow approximately correctedfor leak, calculating an index J varying from 0 to 1 equal to the fuzzyextent to which said respiratory airflow is large positive for longerthan expected, or large negative for longer than expected, calculating amodified airflow equal to said respiratory airflow multiplied by thefuzzy negation of said index J, identifying the phase in the respiratorycycle, calculating the instantaneous inspired volume as the integral ofsaid modified airflow with respect to time, with the integral held atzero during the expiratory portion of the respiratory cycle, calculatinga desired instantaneous delivery pressure as a function at least of thesaid instantaneous inspired volume, and setting delivered pressure tosubject to the desired delivery pressure.
 17. A method for providingventilatory assistance in a spontaneously breathing subject, comprisingthe steps of: determining respiratory airflow approxiamately correctedfor leak, calculating an index J varying from 0 to 1 equal to the fuzzyextent to which the respiratory airflow is large positive for longerthan expected, or large negative for longer than expected, identifyingthe phase in the respiratory cycle, calculating a modified respiratoryairflow equal to the respiratory airflow multiplied by the fuzzynegation of said index J, calculating the instantaneous inspired volumeas the integral of the modified airflow with respect to time, with theintegral held at zero during the expiratory portion of the respiratorycycle, calculating the desired instantaneous delivery pressure as anexpiratory pressure plus a resistance multiplied by the instantaneousrespiratory airflow plus a nonlinear resistance multiplied by therespiratory airflow multiplied by the absolute value of the respiratoryairflow plus an elastance multiplied by the said adjusted instantaneousinspired volume, and setting delivered pressure to subject to thedesired delivery pressure.
 18. A method for providing assistedventilation to match the subject's need, comprising the steps of:describing a desired waveform template function Π(φ), with domain 0 to 1revolution and range 0 to 1, determining respiratory airflowapproximately corrected for leak, calculating an index J varying from 0to 1 equal to the fuzzy extent to which the respiratory airflow is largepositive for longer than expected, or large negative for longer thanexpected, calculating J_(PEAK) equal to the recent peak of the index J,calculating the instantaneous phase in the respiratory cycle,calculating a desired amplitude of pressure modulation, chosen toservo-control the degree of ventilation to at least exceed a specifiedventilation, calculating a desired delivery pressure as an endexpiratory pressure plus the calculated pressure modulation amplitude Amultiplied by the value of the waveform template function Π(φ) at thesaid calculated phase φ, and setting delivered pressure to subject tosaid desired instantaneous delivered pressure.
 19. A method forproviding assisted ventilation to match the subject's need, as claimedin claim 18, in which the step of calculating a desired amplitude ofpressure modulation, chosen to servo-control the degree of ventilationto at least exceed a specified ventilation, comprises the steps of:calculating a target airflow equal to twice the target ventilationdivided by the target respiratory rate, deriving an error term equal tothe absolute value of the instantaneous low pass filtered respiratoryairflow minus the target airflow, and calculating the amplitude ofpressure modulation as the integral of the error term multiplied by again, with the integral clipped to lie between zero and a maximum.
 20. Amethod for providing assisted ventilation to match the subject's need,as claimed in claim 18, in which the step of calculating a desiredamplitude of pressure modulation, chosen to servo-control the degree ofventilation to at least exceed a specified ventilation, comprises thefollowing steps: calculating a target airflow equal to twice the targetventilation divided by the target respiratory rate, deriving an errorterm equal to the absolute value of the instantaneous low pass filteredrespiratory airflow minus the target airflow, calculating an uncorrectedamplitude of pressure modulation as the integral of the error termmultiplied by a gain, with the integral clipped to lie between zero anda maximum, calculating the recent average of said amplitude as the lowpass filtered amplitude, with a time constant of several times thelength of a respiratory cycle, and setting the actual amplitude ofpressure modulation to equal the said low pass filtered amplitudemultiplied by the recent peak jamming index J_(PEAK) plus theuncorrected amplitude multiplied by the fuzzy negation of J_(PEAK). 21.A method for providing assisted ventilation to match the subject's need,and with particular application to subjects with varying respiratorymechanics, insufficient respiratory drive, abnormal chemoreceptorreflexes, hypoventilation syndromes, or Cheyne Stokes breathing,combined with the advantages of proportional assist ventilation adjustedfor sudden changes in leak, comprising the steps, performed at repeatedsampling intervals, of: calculating the instantaneous mask pressure asclaimed in claims 16 or 17, calculating the instantaneous mask pressureas claimed in claim 18, calculating a weighted average of the twopressures, and setting the mask pressure to the said weighted average.